Modern implantable infusion devices, or implantable pumps, for delivering an infusate (e.g., medicaments, insulin, etc.) commonly have a rigid housing that maintains a collapsible infusate reservoir. The housing includes a needle-penetrable septum that covers a reservoir inlet. A flow passage is provided between the reservoir and an exterior surface of the device. At the flow passage outlet, a flexible delivery catheter is provided.
These devices are implanted at a selected location in a patient's body so that (i) the inlet septum is proximate to the patient's skin and (ii) a distal end of the catheter is positioned at a selected delivery site. Infusate can then be delivered to the infusion site by controlling the flow of such fluid from the device infusate reservoir into the delivery catheter. When the infusate reservoir becomes empty, the reservoir is refillable through the reservoir inlet by injecting a new supply of infusate through the apparatus' inlet septum. Due to the location of the device in relation to the skin of the patient, injection can be readily accomplished using a hypodermic needle (or cannula).
Infusate is expelled from the reservoir to an infusion site by collapsing the reservoir. Some infusion pumps use an electrically powered mechanism to “actively” pump infusate from the infusate reservoir into the delivery catheter. Examples of these types of “active pumping” devices include so-called peristaltic pumps (e.g., SynchroMed® implantable pump from Medtronic, Inc., Minneapolis, Minn.) and accumulator-type pumps (e.g., certain external infusion pumps from Minimed, Inc., Northridge, Calif. and Infusaid® implantable pump from Strato/Infusaid, Inc., Norwood, Mass.). These devices have certain advantages; however, such devices have a large disadvantage in that they use relatively large amounts of battery power to effect infusion. Given that batteries tend to add bulk and weight and their replacement requires surgical intervention, it is very desirable to minimize power consumption in implantable infusion pumps.
Another type of implantable pump that is typically much more electrically efficient uses a passive pumping mechanism. In fact, certain of these devices can be constructed and operated without any electrical power at all. A passive pumping mechanism generally consists of a means of pressurizing the infusate reservoir and a means of rstricting the fluid flow. All of these devices operate under the principle that the fluid flow rate (Q) is directly proportional to a pressure difference (ΔP) between the infusate reservoir interior and the delivery site and inversely proportional to the total flow resistance provided by the fluid passage and delivery catheter (collectively (R)), whereinQ=ΔP÷R A practical pump must have a predictable flow rate (Q). To achieve this goal, conventional designs have strived to develop substantially constant pressure sources.
The first means of developing such a “constant pressure source” includes using a two-phase fluid, or propellant, that is contained within the rigid housing and is further confined within a fluid-tight space adjacent to the infusate reservoir. Pumps constructed in this manner are called “gas-driven” pumps.
The propellant is both a liquid and a vapor at patient physiological temperatures, e.g., 98.6° F., and theoretically exerts a positive, constant pressure over a full volume change of the reservoir, thus effecting the delivery of a constant flow of infusate. When the infusate reservoir is expanded upon being refilled, a portion of such vapor reverts to its liquid phase and thereby maintains a state of equilibrium between the fluid and gas states at a “vapor pressure,” which is a characteristic of the propellant. The construction and operation of implantable infusion pumps of this type are described in detail, for example, in U.S. Pat. Nos. 3,731,681 and 3,951,147. Pumps of this type are commercially available, for example Model 3000™ form Arrow International, Reading, Penn. and IsoMed® from Medtronic, Inc., Minneapolis, Minn.
Gas-driven infusion pumps typically provide an electrically efficient means to deliver a flow of infusate throughout a delivery cycle. However, such infusion pumps depend upon a constant pressure source, wherein the output fluid flow rate is directly proportional to a propellant-reservoir pressure. If the propellant-reservoir pressure varies, then so will the fluid flow rate and the drug delivery rate.
The propellant-reservoir pressure of conventional gas-driven infusion pumps are susceptible to changes in ambient temperature and pressure. This, in turn, makes the fluid flow rate such devices susceptible to changes in ambient temperature and pressure. Such changes in drug infusion rates are undesirable and, in certain situations, unacceptable.
Circumstances readily exist where either ambient temperature or pressure can rapidly change a significant amount. For example, the reservoir pressure of some conventional gas-driven pumps can change as much as 0.5 psi for each 1° F. change in body temperature. Thus, for example, assuming a pump driving force of 8 psi at 98.6° F., a fever of only 102.6° F. can result in a twenty-five percent (25%) increase in propellant-reservoir pressure and thus, a corresponding (or larger) increase in an fluid flow rate. In addition, changes in environmental temperature affect the infusate viscosity as well as the vapor pressure produced by the propellant, thereby further increasing the pump's susceptibility to temperature.
An even more serious situation results from changes in ambient pressure. Although minor variations in ambient pressure at any given location on earth may not significantly affect delivery flow rates, with modern modes of transportation, a patient can rapidly change altitude during travel, such as when traveling in the mountains or when traveling by plane. In a like manner, a patient can experience a rapid change in pressure when swimming or diving.
The rigid housing of the conventional, gas-driven infusion pump provides an absolute constant-internal pressure (PR) (at constant temperature) independent of external pressures. However, largely due to compliance by the lungs and venous circulatory system, hydrostatic pressure within the human body closely follows ambient pressure (PD).
The net effect is that the pressure differential (ΔP=PR−PD) in conventional gas-driven pumps changes linearly with ambient pressure. Consequently, a delivered infusate flow rate can increase as much as forty percent (40%) when a patient takes a common commercial airline flight.
To overcome these practical circumstances, some conventional gas-driven infusion devices have been provided with elevated reservoir pressures. The increased reservoir pressures are not intended to prevent variations in a constant pressure delivery but are intended to mitigate their effect. In particular, for any given change of pressure, the effect on flow rate is effectively lessened if the total percentage of pressure change (relative to the reservoir pressure) can be reduced. These infusion devices possess undesirable attributes in that refilling operations are more difficult and the high-pressure vessels, which form the pump housing structures, must necessarily be stronger and are therefore more susceptible to manufacturing problems.
Another method of attempting to produce a constant pressure source and thereby more accurately controlling a rate of fluid delivery is to incorporate a pressure regulator, such as that disclosed in U.S. Pat. No. 4,299,220. The pressure regulator described therein, which is positioned between the infusate reservoir and delivery catheter, uses a diaphragm valve to maintain a constant pressure differential (OP) across the fluid flow restrictor. In addition to increasing the operational complexity of such a pump mechanism and the volume of a fluid path extending therethrough, the pressure regulator may, depending upon device configuration, subject infusate solutions to high local shear stresses, which may alter the chemical or therapeutic properties of certain infusates.
An alternative method for attempting to produce a constant pressure source and thereby more accurately controlling a rate of fluid delivery, as well as addressing the susceptibilities of the two-phase pumps to ambient temperature and pressure, is proposed in U.S. Pat. No. 4,772,263. Specifically, in place of the conventional rigid enclosure that maintains a two-phase fluid, the disclosure teaches forming the fluid reservoir between a rigid portion (which maintains at least the inlet septum) and a flexible drive-spring diaphragm. The spring diaphragm is exposed to the body of the patient and the pressure therein. The spring diaphragm creates a more desirable “relative” (as opposed to an absolute) pressure source. By exposing the spring diaphragm to the pressures inside the body it is possible for the pump to respond and react to changes in ambient pressure so that AP is unaffected. Likewise, it is possible to construct the spring diaphragm so that the pressure that it generates is not affected by changes in ambient temperature. While this configuration offers practical performance advantages, this design offers a unique configuration that may not be adopted by all constant flow pump designs.
Accordingly, a need exists for an electrically efficient system to enable a controllable (constant or variable) fluid flow delivery rate independent of either a constant reservoir pressure or external conditions that may otherwise result in undesirable or unpredictable output fluid flow variations.